Methods, apparatuses and systems for amputee gait capacity assessment

ABSTRACT

The present invention includes a tethered ankle-foot prosthesis emulator having two independently-actuated forefoot digits that are coordinated to provide plantarflexion and inversion-eversion torques, and an independently-actuated heel. An end-effector was designed which is worn by a subject, and which was integrated with off-board motor and control hardware, to facilitate high bandwidth torque control. The platform is suitable for haptic rendering of virtual devices in experiments with humans, which may reveal strategies for improving balance or allow controlled comparisons of conventional prosthesis features. A similar morphology is also effective for autonomous devices.

RELATED APPLICATIONS

This application is a continuation-in-part of pending U.S. patentapplication Ser. No. 14/827,299, filed Aug. 15, 2015, which claims thebenefit of U.S. Provisional Application Ser. No. 62/070,134, filed Aug.15, 2014 and U.S. Provisional Application Ser. No. 62/230,046, filed May26, 2015.

STATEMENT REGARDING GOVERNMENT-SPONSORED RESEARCH AND DEVELOPMENT

This invention was made with partial government support under NIH grant1R43HD076518-0, NSF grant CMMI-1300804 and NSF grant CBET1511177. Thegovernment has certain rights in this invention.

FIELD OF THE INVENTION

This invention relates to methods, apparatuses and systems for assessinggait capacity of lower-limb amputees, as an aid in prescribingappropriate prosthetic devices. In particular, the invention relates torobotic emulation of commercially-available prostheses.

BACKGROUND OF THE INVENTION

Over one million people in the U.S. live with limb loss, with anestimated 100,000 new cases each year, over 80% involving the lowerlimb. Commercial prostheses are available with many different designsand features, at prices ranging from a few hundred dollars to over onehundred thousand dollars. Yet the prescription process is based mostlyon subjective assessments and past performance, with no way toprospectively determine whether an increment in cost will yield asatisfactory improvement in quality of life for a given individual.

The most important mobility issues are discomfort, stability andfatigue. Persons with below-knee amputation choose a lower self-selectedwalking speed than able-bodied persons, and expend at least 20% moreenergy to walk at the same speed. Many advanced conventional prostheseshave features such as higher elastic energy storage and return, andusers generally prefer these feet for reasons of comfort. Nonetheless,the speed and energy cost discrepancies from non-amputees persistdespite a wide range of complexity and cost in commercial prostheticfeet.

Recently, advanced foot prostheses have come to market promising tobreak this barrier by restoring one crucial component all passiveprostheses lack: an ankle joint that can perform net positive work onthe body. One such device (the BiOM T2 System) can improve walkingmechanics, returning walking speed and energy expenditure to near-normallevels for some patients. It is unclear, however, whether allindividuals will benefit, and possible benefits come at a steep cost:such devices are currently considerably more expensive, accessible toonly the best-reimbursed or wealthiest patients. This would be a badinvestment for any patient who does not realize major gains in qualityof life, especially since advanced devices often cannot be returned orresold following initial use. Emerging robotic prostheses like the BiOMintensify a longstanding dilemma in prosthetics practice: how canpractitioners and insurance companies identify who will benefitsufficiently from increased performance to justify the higher cost ofadvanced devices? This problem has recently become more acute, asMedicare and other payers have identified cases of fraud, and inresponse have increased the demand for documentation to supportclassification of each individual's gait capacity. The argument is thatcurrent practice for assigning a K-level (KO to K4) relies too heavilyon unreliable information such as prosthetist's opinion and thepatient's stated activities and goals, and so can be manipulated, to thedetriment of the payer.

Recent advances have added some nuance to the differentiation amongK-levels, such as short in-clinic functional mobility tests orapproximate activity classification based on time-binned stepclustering. These tests include tasks such as freely-selected walking,standing and sitting transitions, climbing and descending stairs,navigating obstacles, and single-limb standing. However, all of thesecategorization methods have a common limitation: they are based oncurrent mobility with the patient's current conventional prosthesis.They do not incorporate any information on how an individual patientwill use and respond to a more advanced prosthesis, such as the BiOM T2.There is essentially no information available to help clinicians andpayers determine whether a particular patient will benefit from anadvanced prosthesis. There is therefore a high probability of suboptimalpatient outcomes, economic inefficiency, and provider-carrier conflictduring the prescription of advanced prostheses.

Robotic prostheses can improve locomotor performance for individuals whohave restricted mobility due to lower-limb amputation. During walking,these devices can restore normal ankle and knee kinematics, reducemetabolic rate, and provide direct neural control of the limb. Asrobotic technologies improve, active prostheses are expected to enhanceperformance even further.

Ankle inversion-eversion, or roll, is an important aspect of prosthesisfunction. Commercial prostheses typically include a passiveinversion-eversion degree of freedom, either using an explicit joint ora flexure. This mitigates undesirable torques created by uneven ground.Inversion torque has a strong effect on side-to-side motions of the bodyduring human walking, and its pattern is altered among individuals withamputation. Side-to-side motions seem to be less stable in bipedallocomotion, particularly for amputees. Difficulty controllinginversion-eversion torque in the prosthetic ankle may partially explainreduced stability and increased fear of falling and fall rates amongpeople with amputation.

Robotic prosthesis designs have begun to incorporate active control ofankle inversion-eversion. A tethered ankle prosthesis with inversionprovided by a four-bar linkage and controlled by a linear actuator hasbeen described, in which a plantarflexion degree of freedom is providedusing a passive spring. A prototype device intended to provide bothplantarflexion and inversion-eversion control using two motors and agimbal joint has also been described.

The mass of prostheses with active inversion-eversion control isgenerally related to joint design. Linkages and gimbal joints ofteninvolve large parts with complex loading, resulting in increasedstrength and mass requirements. An alternative is suggested by thesplit-toe flexures in conventional passive prostheses and the actuationschemes in some powered ankle orthoses. During walking, peakinversion-eversion torques are of much lower magnitude than peakplantarflexion torques, and the majority of the inversion impulse occursduring periods of high plantarflexion torque. Coupling plantarflexionand inversion-eversion torque through the actions of two hinged forefootdigits might provide sufficient torque capacity, enabling a simple,lightweight design.

Mechatronic performance in experimental prosthesis systems can also beimproved by separating actuation hardware from worn elements. A tetheredemulator approach decouples the problems of discovering desirableprosthesis functionality from the challenges of developing fullyautonomous systems. Powerful off-board motors and controllers can beconnected to lightweight instrumented end-effectors via flexibletethers, resulting in low worn mass and high-fidelity torque control.Such systems can be used to haptically render virtual prostheses tohuman users, facilitating the discovery of novel device behaviors thatcan then be embedded in separate autonomous designs. This approach canalso be used for rapid comparison of commercial prostheses in a clinicalsetting. To be most effective, such prosthesis emulators should havehigh closed-loop torque bandwidth and lightweight, strong,accurately-instrumented end-effectors.

Torque control in robotic emulator systems can be improved withappropriate series elasticity. Adding a spring in series with ahigh-stiffness transmission can reduce sensitivity to unexpectedactuator displacements imposed by the human. Unfortunately, thiscompliance also reduces force bandwidth when the output is fixed,because the motor must displace further when stretching the spring. In atethered system, the flexible transmission itself is likely to havesignificant compliance, which might provide appropriate serieselasticity.

Prosthesis emulators have not yet explored the full range of controlstrategies. For example, humans appear to control center of pressureduring walking or maintain similar roll-over shapes despite varyingwalking conditions. To study these control strategies, an emulator mustbe equipped with at least two degrees of freedom in both frontal andsagittal planes.

SUMMARY OF THE INVENTION

The present invention describes methods, apparatuses and systems forassessing gait capacity of lower-limb amputees, as an aid in prescribingappropriate prosthetic devices. The invention describes a system toemulate the characteristics of various types of prosthetic devices toprovide an amputee the opportunity to simulate the use of such variousprosthetic devices. The system collects and analyzes a variety of dataduring the simulated use of prosthetic devices to provide quantitativeinformation on the appropriateness of various prosthetic devices for theindividual amputee.

The platform is suitable for haptic rendering of virtual devices inexperiments with humans, which may reveal strategies for improvingbalance or allow controlled comparisons of conventional prosthesisfeatures. A similar morphology may be effective for autonomous devices.

In one embodiment, the present invention includes a tethered ankle-footprosthesis with two independently-actuated forefoot digits that arecoordinated to provide both plantarflexion and inversion-eversiontorques. This configuration allows a simple lightweight structure. Thetwo digits extend the capabilities of the original invention toemulation of so-called multi-axial or split-toe prosthetic feet.

In a second embodiment, the present invention also includes an actuatedheel. Actuation of the three digits allows the adjustment of the originand magnitude of the ground reaction force vector. This additionaldegree of control extends the capability of the original invention tomore accurate emulations, for example, emulation of prosthetic footrollover shape. The platform is suitable for haptic rendering of virtualdevices in experiments with humans, which may reveal strategies forimproving balance or allow controlled comparisons of conventionalprosthesis features. A similar morphology may be effective forautonomous devices.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows the apparatuses of a first embodiment of the presentinvention in schematic form;

FIG. 2 shows use of the device of the present invention by a subject;

FIG. 3 shows results of benchtop testing of an exemplar embodiment ofthe present invention;

FIG. 4 shows results of testing of an exemplar embodiment of the presentinvention with test subjects, in particular, FIG. 4 showsimpedance-matching performance of ankle movement vs. angle in threeemulator modes and one custom mode;

FIG. 5 shows impedance matching design for emulation of a normal ankle;

FIG. 6 illustrates a second embodiment of the invention utilizing amechanical design of the two degree of freedom ankle-foot prosthesisemulator;

FIG. 7 shows coupling between prosthesis plantarflexion andinversion-eversion torque illustrated with typical human walking data;

FIG. 8 presents benchtop tests demonstrate low torque measurement error.high peak torque and high closed-loop torque bandwidth in bothplantarflexion and inversion-eversion directions;

FIG. 9 shows a graph depicting disturbance rejection;

FIG. 10 presents graphs depicting torque tracking during walkingexperiments.

FIG. 11 shows the detail of the construction and components of theend-effector for the three-digit embodiment of the emulator.

FIG. 12 shows an overall system diagram of the three degree of freedomembodiment of the emulator, including the controller, motors andend-effector.

FIG. 13 shows the three degrees of freedom of movement of the ankle-footembodiment of the emulator, including two forefoot digits and a heeldigit.

FIG. 14 shows the three modes of movement of the three degree of freedomembodiment of the emulator.

FIG. 15 shows the support triangle formed by the contact points of theheel and two forefoot digits in the three degree of freedom embodimentof the emulator.

DETAILED DESCRIPTION OF THE INVENTION

For purposes of this invention, the term digit shall refer genericallyto a forefoot digit, a heel, or any other portion of the end-effectorcontrolled by a motor controlled by one or more control components.

This invention describes methods, apparatuses and systems for assessinggait capacity of lower-limb amputees, as an aid in prescribingappropriate prosthetic devices. The invention describes a system toemulate the characteristics of various types of prosthetic devices toprovide an amputee the opportunity to test-drive various prostheticdevices. The system collects and analyzes a variety of data during thesimulated use of prosthetic devices to provide quantitative informationon the appropriateness of various prosthetic devices for the individualamputee.

The apparatus is an off-board, actuated, powered ankle-foot prosthesisemulator that attaches to the tibial pylon of an amputee's prostheticlower limb in place of their prescribed prosthetic foot (FIG. 1). Theemulator provides programmable, realtime-controlled torque about theankle, suitable for emulating different classes of commercial prosthesisranging from dissipative block-and-foam prostheses (e.g. Solid AnkleCushion Heal, or “SACH”), to resilient dynamic elastic response designs(e.g. Dynamic Elastic Response or “DER”, Elastic Storage and Return, or“ESR” or “ESAR”; for example, Flexfoot), to dynamic elastic feet withmulti-axial capabilities, to prostheses that automatically adjust ankleangle (e.g. OssurProprio, Endolite élan, OrthoCare, InnovationsMagellan), to active, powered devices (e.g. BiOM T2 System, SPARKy andFreedom Innovations Foot). The emulator allows an amputee to experiencethe advantages and disadvantages of each level of technology in rapidsuccession. This system provides objective data about the capacity of anindividual patient for gait improvement, and thereby helps practitionersdetermine the prescription that best balances mobility gains andfinancial costs.

The emulation programs are based on published curves of ankle torque vs.ankle angle over a complete gait cycle. The differences between deviceclasses include the ankle's angular quasi-stiffness and the net workperformed during a gait cycle. Net prosthesis work is an importantdifferentiating factor among these classes; SACH feet provide about 55%energy return, DER about 80%, and powered robotic about 120% energyreturn, i.e. it supplies positive net work. Example emulationperformance (see FIG. 5), demonstrates excellent emulation of SACH andDER modes and reproduction of characteristic positive work output forthe BiOM.

The emulation system enables the characterization of responses toconditions that both exceed and under-perform currently availablecommercial prostheses. In cases where functional benefits may be unclearfrom comparisons of settings emulating conventional, DER and poweredprostheses, it may be valuable to test the individual's response toexcessive energy gain or loss in the prosthesis. Patients who readilyadapt to unusual conditions may be better candidates for prostheses withcomplex control behaviors. Clinical feedback to manufacturers could alsospur development of products that are better suited to their clients.Because these cases cannot be achieved by commercial devices, such testscan only be done with an emulator such as described herein.

This invention enables evidence-based medicine in prosthetics. Theavailability of objective gait capacity data closes the feedback loopbetween the prosthetics clinic and the physician. This invention enablesa work flow that starts in the prosthetics clinic with the emulatorsystem. Test results for metrics such as effort, gait stability, maximumspeed, and user preference with different classes of prosthesis willbecome part of the patient's medical record, accessible to thephysician, the prosthetist and the payer. The physician will considerthis information about gait capacity alongside other medicalindications, and make a better-informed, more appropriate, and moredefensible prosthesis prescription. The prosthetist will use testresults to choose specific componentry and tune the prosthesis to meetthe patient's documented needs. For example, a subject who achieves highspeeds with moderate effort, but demonstrates lateral instability, maybe given a DER foot with an especially wide base of support. Thisinvention provides quantitative evidence to these decisions, which wouldotherwise be made by subjective visual assessment.

A first embodiment of the present invention is shown in FIG. 1. Shown inFIG. 1C, is end-effector 100 which is worn by subject 10. End-effector100 connects to the subject using standard universal adapter 124. Acustom-sized spacer 128, shown in FIG. 2, connects to universal adapter124 and allows connection to prosthesis 12 worn by subject 10.

Forefoot portion 102 of end-effector 100 pivots around ankle joint 104.The tension on forefoot 102 is controlled by a series of leaf springs108 which are connected via chain 127 to sprocket 116. The tensionexerted through leaf springs 108 is controlled by transmission element134, as shown in FIG. 1, which is attached via transmission toattachment 112 to pulley 114. Dorsiflexion spring 126 ensures thatforefoot 102 returns to its un-flexed position after tension has beenreleased by chain 127. The angle of the ankle displacement is sent byankle encoder 106 and is sent back to prosthesis control components 132and 136 via tether 134 b.

Heel 111 of end-effector 100 is connected to the ankle joint 104 viapassive heel spring 110. Pulley encoder 118 encodes the rotation ofpulley 114 and sends that information to prosthesis control components132 and 136 via tether 134 b.

Off-board motor 130 as shown in FIG. 2 controls end-effector 100 underclosed-loop computer control. Flexible tether 134 b provides feedbackfrom sensors mounted on end-effector 100 and transmission cable 134 acontrols the tension exerted by pulley 114. Using off-board motor 130and control components 132 and 136 allows a more flexible, muchhigher-performance system with a simpler design and less body-mountedmass than an untethered system. With this design, the mechatronicperformance of the prosthesis is dominated not by the mechanicalproperties of end-effector 100, but instead by the closed-loop actuationspecified under computer control. Thus, a single ankle-foot end-effector100 can emulate many control behaviors and mechanical elements. In apreferred embodiment, prosthesis control components 132 and 136 containmodels of a plurality of commercially-available prosthesis and is ableto simulate the feel of these prostheses to subject 10.

In an exemplar embodiment, the emulator system comprises a 1.6 kWservomotor and motor controller (Baldor Electric Co.), a 1 GHz controlhardware package (dSPACE, Inc.) to perform real-time torque commandtracking, and a 3 m long Bowden cable (steel and polyethylene sheath and6 mm Vectran drive cable, similar to a bicycle brake cable) to transferpower to end-effector 100. The prosthesis end-effector uses a fiberglassleaf spring 108 for series elasticity while transforming cable tensioninto a torque on the forefoot segment 102, equivalent to ankleplantarflexion in the human foot, as shown in FIG. 1B. Ankle jointangles are measured directly with sensors or encoders 106, while jointtorque is inferred from spring deflection based on calibration results.As shown in FIG. 3, the heel 111 comprises a passive leaf spring 108,which provides cushioning at load acceptance. Benchtop testing of theexemplar system (See FIG. 4) established a 17 Hz bandwidth for torqueresponse, 95% rise and fall time of less than 0.070 s, peak power outputof 1060 W and peak ankle torque of 175 N·m, with worn mass less than 1kg. This combination of low worn mass and high closed-loop torquebandwidth are the keys to high-fidelity emulation of specializedprosthetic devices. See Table 1.

Joint Torque Torque Bandwidth Worn Power Control System [N · m] [Hz]Mass [kg] [peak, W] Tethered Joint Style Michigan Pros.² 180 2 2.8 360Yes Ankle Pros. BLEEX^(3,4)  175* n.r. 2.50 200 No Ankle Exo. VanderbiltPros.⁵ 130 n.r. 2.50† 250 No Ankle Pros. MIT Pros.^(6,7) 120 3.8 2.50350 No Ankle Pros. Michigan Exo.⁸ 120 2 1.37 n.r. Yes Ankle Exo. MITAFO⁹ 120 n.r. 1.12 370 Backpack Ankle Exo. SPARKY^(10,11) 100 n.r. 2.70400 No Ankle Pros. MINDWALKER¹² 100 6 2.90 960 Yes n/a Exo.Anklebot^(13,14) 50 n.r. 3.60 340 Yes Ankle Exo. RoboKnee^(15,16) 40 7.53.00 630 No Knee Exo.

Table 1: Comparison of candidate joint torque control systems capable ofankle-like torques (at least 40 N·m, or ⅓ the typically-observed valuesfor normal human walking). Peak torque, closed-loop torque bandwidth (−3dB gain), total mass worn by the human, and peak joint mechanical powerare compared to experimental data from the emulator of the presentinvention. The inventive system is stronger, lighter, more responsive,and more powerful than reported values for all prior systems. Bestperformance is indicated in bold. The abbreviations n/a and n.r.indicate not applicable and not reported, respectively.

The invention includes methods and systems to distinguish the gaitcapacity of patients to be fitted with prosthetics. An exemplar systemas described above was tested over a range of work outputs on fiveunilateral amputee subjects. In this experiment, net ankle push-off workwas varied from −1.9 to 8.0 times the net work generated by the healthybiological ankle during normal walking, as shown in FIG. 4. For twosubjects, metabolic rate was minimized near the level of work providedby the healthy biological ankle; for two subjects, metabolic rate wasunaffected by push-off work; and for one subject, increased push-offwork increased metabolic rate. These very different effects show theneed for device selection based on each individual's response.

In addition to being able to test for different levels of work, theinvention incorporates specific emulations of commercial prostheses. Toperform such emulations, the invention uses impedance-matching control,in which the ankle's angular stiffness (N·m per radian) is controlled tomatch a reference value determined from a commercial prosthesis of eachclass. In the demonstration of the exemplar system, published anklekinetics data was used to build a reference curve of ankle moment vs.ankle angle for examples of a SACH foot, a DER foot, and the BiOM T2System foot.

Those skilled in the art will recognize that the present invention couldbe adapted to include such data from other existing and futureprosthetic devices. Also, the system can be extended through expressionof finer and finer device categories, ultimately including specificparameters of a given make and model of prosthesis. In such a case, itmay be beneficial to have a second assessment in which the prescriptionis refined to specific values of, e.g., keel or bumper stiffness anddamping. Another extension is to add automatic ankle angle adjustment,to identify any advantages to the individual of prostheses that adapt toterrain (e.g. Proprio, élan, Magellan). When additional fully roboticprostheses come to market, realistic emulations of each could beincluded in the system, so as to identify which specific device suits anindividual. The invention could be extended to includemulti-degree-of-freedom emulators that provide additional control, forexample over center of pressure in the frontal plane

The target for the impedance-matching controller was designed using apiecewise-linear approximation to the nonlinear reference curve shown inFIG. 5. The ground contact period was divided into four sub-phases:early and late dorsiflexion and early and late plantarflexion. Eachsub-phase was given a separate slope (effective stiffness) andintercept. Transitions from one sub-phase to the next were triggered bythe direction of ankle velocity (dorsiflexing or plantarflexing) and bythreshold values of ankle angle (early or late). Using differentstiffness targets in each sub-phase allowed the reference curve toapproximate the stiffness and energy return (or generation)characteristics for each commercial prosthesis. In addition to theseemulation modes, we designed a high power condition (HiPow) withsignificantly greater work output than the natural ankle provides. Thismode was designed with a dorsiflexion profile typical of non-amputeewalking, but very low stiffness during early plantarflexion (shallowslope), such that the ankle provided high moment over significantdeflection, resulting in positive ankle work roughly twice that of thenatural ankle. This result is shown in FIG. 4D.

The impedance-matching emulation approach described above yielded goodbehavior in reproducing target work output, in addition to reproducingmoment vs. angle reference behavior from the commercial prostheses, asshown in FIG. 4. Emulations of SACH, DER and BiOM feet were each veryclose to the reference data (moment vs. time RMS error of 7 N·m, 2 N·mand 4 N·m, respectively). The HiPow mode successfully produced a veryhigh work output and gave users an ambulation experience they could nothave had with any other system, though its trajectory trackingperformance was slightly worse (RMS error of 8 N·m from its designtarget). To improve this error, in one embodiment of the invention,iterative learning control is used to reduce the residual errors intarget tracking.

In one aspect of the present invention, the system collects data onmetabolic rate and/or heart rate of the subject 10 using respirometerand pulse oximeter 20, shown being worn by subject 10 in FIG. 2, whilesubject 10 is using the emulator on a flat or tilted treadmill 14, toestimate maximum sustainable speed.

In one aspect of the present invention, the system collects data onmaximum walking speed, for use in estimating gait capacity.

In one aspect of the present invention, the system includes inertialsensor on the emulator and the contralateral foot to enable assessmentof gait stability using variability in stride length, stride width andstride time. Variable cadence is important to high K-level mobilityratings, where greater cadence variability is considered bettermobility. At the same time, high kinematic variability within asteady-state task is sometimes associated with fall risk and poormobility. The assessment protocol of the present invention providestests of walking ability at different speeds (and hence cadences), andthe stride variability data measures gait quality during each condition.This data helps to identify more and less beneficial prosthetic devicesfor each individual.

In one aspect of the present invention, data is collected on approximateground reaction force (“GRF”) peaks on the prosthetic side foot. GRFoutcomes can thereby be reported, such as early and late force peaks,without the need for an instrumented treadmill. GRF peaks are importantindicators of gait function such as weight acceptance mechanics and latestance weight support. Comparing force profiles across conditionsprovides additional evidence for appropriate prosthetic prescription.

In a clinical setting, the use of the present invention begins with anassessment of various emulated prosthetics in a prosthetics clinic or ina hospital. The report, including metrics for effort, maximumsustainable speed, gait stability, ground force peaks, and user andassessor feedback on each condition, will be sent to the physician andthe prosthetist. The physician will consider the emulator resultsalongside traditional indications such as general health, desiredactivities, specific residual limb properties, balance confidence, andcost. The addition of objective performance results will allow thephysician to make better-informed, more appropriate, and more defensibleprescriptions. Then, the prosthetist will use the test results to choosespecific componentry and tune the prosthesis to meet the patient'sdocumented needs. For example, a subject who achieves high speeds withmoderate effort in the DER condition, but demonstrates lateralinstability, may be given a DER foot with an especially wide base ofsupport. Or, a subject who can walk twice as fast with the BiOMemulation could be prescribed one based on the large benefit itprovides. As another example, a subject whose speeds and efforts aresimilar with the DER and BiOM emulations could be prescribed ahigh-quality DER foot, because they may gain more from its durabilitythan from the BiOM's positive work output.

A second embodiment of the invention is shown in FIG. 6, having alateral forefoot 202 and a medial forefoot 204 which are independentlycontrolled and sensed. The design and evaluation of the roboticankle-foot prosthesis emulator system has active control of bothplantarflexion and inversion-eversion torques and allowsinversion-eversion using two articulated forefoot digits. End-effector200 was integrated with existing off-board motor and control hardware tofacilitate high bandwidth torque control. End-effector 200 did notinclude explicit series elasticity, testing the sufficiency of axialcompliance in the tether. A basic walking controller was implemented,intended to evaluate the system's potential for emulating prosthesisbehavior during interactions with a human user. This approach wasexpected to result in validation of a system that can explore newdimensions of prosthesis assistance, particularly those related tobalance during walking.

The emulator system consists of off-board actuation and control hardware214, a flexible Bowden-cable tether 210 and end-effector 200 worn bysubject 10. The prosthesis end-effector 200 has twoindependently-actuated forefoot digits 202, 204 and a separate, passiveheel spring 209. Plantarflexion occurs when both digits 202, 204 rotatetogether, and inversion/eversion occurs when the medial forefoot 204 andlateral forefoot 202 move in opposite directions. Plantarflexion andinversion-eversion torques are proportional to the sum and difference,respectively, of individual forefoot torques. The prototype used inexperiments is instrumented with encoders 211 a and 211 b at each anklejoint and four strain gages 213 in a Wheatstone bridge on each forefootto measure torque. End-effector 200 is connected to the user via auniversal pyramidal adapter 215. Rubber bands 217 dorsiflex the forefootdigits during the swing phase of walking.

End-effector 200 was designed and constructed with torque control inboth plantarflexion and inversion-eversion directions. As with the firstembodiment, the actuation and control hardware is located off-board soas to keep worn mass low. Flexible Bowden-cable tethers 210 transmitmechanical power to the prosthesis but do not interfere with naturalmovements of the limb.

End-effector 200 achieves torque and motion in both plantarflexion andinversion-eversion directions using two independent forefoot digits 202and 204. The digits share a single axis of rotation 212 similar to theplantarflexion axis in the human ankle joint, and are spacedmediolaterally such that one is closer to the centerline of the body, asshown in FIG. 6B. Plantarflexion occurs when both digits rotate in thesame direction, and inversion-eversion occurs when the digits rotate inopposite directions, as shown in FIG. 6C. For purposes of thisembodiment of the invention, plantarflexion angle is defined as theaverage of the forefoot angles and the inversion-eversion angle as thedifference between the forefoot angles multiplied by the ratio offorefoot length to half the foot width. Similarly, plantarflexiontorque, τ_(ρl), is defined as the sum of the lateral and medial forefoottorques, τ_(l) and τ_(m), while inversion torque, τ_(inv), is defined asthe difference between the lateral and medial forefoot torquesmultiplied by the ratio of half the foot width, ½ w, to forefoot length,l, or

$\begin{matrix}{\tau_{\rho\; l} = {\tau_{l} + \tau_{m}}} & (1) \\{\tau_{inv} = {\frac{w}{2\; l}\left( {\tau_{l} - \tau_{m}} \right)}} & \;\end{matrix}$

The end-effector 200 consists of a frame, two forefoot digits withrevolute joints, and a compliant heel 209. The frame of the device isconnected to the pylon or socket of subject 10 via universal pyramidaladapter 215. The frame houses needle roller bearings for ankle joints,which have a double-shear construction. Each forefoot is long and thin,tapers towards its ends, and has an I-beam cross section, making itwell-suited to three-point bending. One end of the forefoot contacts theground, while the other end is acted on by cable 219 (only one shown inFIG. 6), with the hinge located in the middle. When cable 219 pullsupward on the posterior aspect of the forefoot, a moment is generated.The conduit of cable 219 presses down on the frame equally andoppositely, such that the foot experiences no net force from thetransmission. Rubber bands 217 act to dorsiflex the forefoot when thetransmission allows, such as during the swing phase. A separate,unactuated heel spring 209 is connected to the frame. Rubber-coatedplastic pads are attached to the ends of the heel and forefoot digitsfor better ground contact. The frame and forefoot digits were machinedfrom 7075-T6 aluminum, the heel spring was machined from fiberglass(GC-67-UB, Gordon Composites), and the forefoot pads were fabricatedusing fused-deposition modeling of ABS plastic. The cable is preferablya Bowden-style cable.

The dimensions of end-effector 200 were based on an average male humanfoot. In the test embodiment, the device measures 0.23 m in length, heelto forefoot tip, 0.07 m in width, forefoot center to forefoot center,and 0.08 m in height, from ground to ankle joint. The forefoot length,from axis of rotation to tip, is 0.14 m. Ankle range of motion is −20 to30 in plantarflexion and greater than −20 to 20 in inversion-eversion.End-effector 200 weighs approximately 0.72 kg.

Medial and lateral forefoot joint angles were sensed individually usingdigital absolute magnetic encoders (MAE3, US Digital). Forefoot torqueswere sensed using strain gages 213 (SGD-3, Omega Engineering),configured in a Wheatstone bridge, with two gages on the top and bottomsurfaces of each forefoot midway between the tip and the ankle joint.Strain gauges 213 measure strain in each rotating forefoot component. Acalibration is then performed to map strain gauge measurements to ankletorque (they are related linearly by the stiffness of the forefootcomponent).

Bridge voltage was amplified (FSH01449, Futek Advanced SensorTechnology, Inc.), sampled at a frequency of 5000 Hz and low-passfiltered with a cutoff frequency of 100 Hz. Plantarflexion andinversion/eversion angles and torques were calculated in software frommedial and lateral forefoot values.

Forefoot digits 202, 204 are actuated through independent Bowden cabletethers 210 and off-board motors, allowing independent control of medialforefoot 204 and lateral forefoot 202 torques. Plantarflexion andinversion-eversion torques can be independently controlled, but maximumallowable inversion-eversion torque is proportional to plantarflexiontorque. When inversion-eversion torque is zero, the plantarflexiontorque is divided evenly between the forefoot digits. As inversiontorque increases towards its limit, the torque on lateral forefoot 202approaches the total desired plantarflexion torque, while the torque onmedial forefoot 204 approaches zero. When inversion (or eversion) torqueequals plantarflexion torque divided by 2l/w, the inversion-eversiontorque cannot be increased further, as doing so would require negativetorque on the medial (or lateral) forefoot, and negative ground reactionforces. This defines a feasible region of inversion torques as afunction of plantarflexion torque (FIG. 7).

FIG. 7 shows coupling between prosthesis plantarflexion andinversion-eversion torque illustrated with typical human walking data.Maximum feasible inversion/eversion torque (gray region) is proportionalto plantarflexion torque (Eq. 1). With a typical plantarflexion torquepattern (solid line) the typical inversion-eversion torque (dashed line)falls within the feasible region for this device. Reference data forhuman walking at 1.6 m/s were used.

Classical feedback control was used to regulate torque during benchtoptests, with an additional iterative learning term during walking trials.Desired torque for each forefoot was first calculated from desiredplantarflexion and inversion-eversion torques. Motor velocities werethen commanded using proportional control on forefoot torque error.Motor velocity is similar to the rate of change in forefoot torque,owing to compliance between the off-board motor and prosthesis forefoot.During walking trials, an additional time-based iterative learning termwas added, which provided feed-forward compensation of torque errorsthat tended to occur at the same time each step.

In walking trials, torque control was used during stance and positioncontrol was used during swing. Initial forefoot contact was sensed froman increase in forefoot torque upon making contact with the ground.During the ensuing stance period, desired inversion-eversion torque wasset to a constant value, providing a simple demonstration of platformcapabilities.

Desired plantarflexion torque during stance was calculated as a functionof plantarflexion angle so as to approximate the torque-anglerelationship observed during normal walking. The end of the groundcontact was detected when plantarflexion torque crossed a minimumthreshold. During the ensuing swing phase, the forefoot digits wereposition controlled to provide ground clearance.

Benchtop tests were conducted to characterize device performance interms of torque measurement accuracy, response time, bandwidth, anddisturbance rejection. Walking trials were performed to assessmechatronic performance under similar conditions as expected duringbiomechanics experiments.

Torque measurement calibration was performed by applying known forces tothe end of each forefoot using free weights and fitting amplified straingage bridge voltage to applied torque. Measurement accuracy wascharacterized in a separate validation test as root mean squared (RMS)error between applied and measured forefoot torques.

Step response tests were performed in which the prosthesis frame andforefoot digits were rigidly fixed and commanded desired torque as asquare wave from 0 to 180 N·m in plantarflexion or −20 to 20 N·m ininversion/eversion. Trials were conducted for each direction andcomputed the mean and standard deviation of the 90% rise and fall times.

Bandwidth tests were performed in which desired torque was commanded asa 0 to 40 Hz chirp, oscillating between 10 and 90 N·m for plantarflexionand between −20 and 20 N·m for inversion-eversion. An exponential chirpwas used to improve signal to noise ratio in the low frequency range.The desired and measured torque was transformed into the frequencydomain using a Fast Fourier Transform and the magnitude ratio and phasedifference was used to generate a Bode plot. The gain-limited andphase-limited bandwidths were calculated as the frequencies at which theamplitude ratio was −3 dB and the phase margin was 45°, respectively.Trials were performed for both torques and calculated crossoverfrequency means and standard deviations.

A test intended to evaluate the torque errors that would arise fromunexpected disturbances to forefoot position was also performed. It wasexpected that high series stiffness in this system might have providedhigh bandwidth at the cost of higher sensitivity to positiondisturbances, for example during initial forefoot contact with theground. The forefoot digits were placed on opposite ends of aseesaw-like testing jig such that forefoot forces were equal andforefoot motions were equal and opposite. A 0 to 25 Hz chirp was appliedin medial forefoot position, oscillating between 0° and 5° ofplantarflexion (or 0 and 0.012 m of forefoot tip displacement) andcommanded a constant desired torque of 30 N·m to the lateral forefoot.The amplitude of the resulting torque error was transformed into thefrequency domain using a Fast Fourier Transform, reported as a percentof the constant desired torque magnitude. The frequency at which errorrose above 30% of the desired torque was calculated, analogous to the −3dB (70% amplitude) criteria used in bandwidth tests.

Walking trials were performed to evaluate torque tracking performanceunder realistic conditions. One subject (67 kg, 1.77 m tall, 23 yrs,male) without amputation wore the device using a simulator boot. Fivewalking trials were conducted in which desired inversion/eversiontorque, τ_(inv) was commanded as: Maximum, 15 N·m, 0 N·m, −15 N·m, andMaximum Negative. The magnitudes of Maximum and Maximum Negativeinversion-eversion torque were proportional to plantarflexion torque ateach instant in time. The subject walked on a treadmill at 1.25 m/s for100 strides in each condition. Each step was normalized to percentstance period and calculated an average step for each condition. Torquetracking error was characterized as both the RMS error across the entiretrial and as the RMS error of the average step. Human biomechanicalresponse was not measured, since this study was intended to evaluateperformance of the robotic system and not the effects of a proposedintervention.

FIG. 8 shows results of benchtop tests to demonstrate low torquemeasurement error, high peak torque and high closed-loop torquebandwidth in both plantarflexion and inversion-eversion directions. Peakplantarflexion torque was at least 180 N·m, and inversion-eversiontorque had a range of at least 20 to 20 N·m. Rise and fall times rangedfrom 0.024 to 0.033 s. Bode plots for closed-loop control of Eplantarflexion and F inversion-eversion torque, calculated fromresponses to 90 N·m and ±20 N·m magnitude chirps in desired torque,respectively. Bandwidth ranged from 20 to 30 Hz, limited by 45° phasemargin.

The benchtop tests revealed low torque measurement error, high peaktorque and high closed-loop torque bandwidth. The root mean squared(RMS) torque measurement errors for medial and lateral forefoot digitswere 1.64 N·m and 2.43 N·m, respectively, following calibration (FIGS.8A & 9B). The 90% rise and fall times between 0 and 180 N·m inplantarflexion torque were 0.033±0.001 s and 0.024±0.001 s (mean±s.d.),with 0.5% and 1.6% overshoot, respectively, as shown in FIG. 9C. The 90%rise and fall times between −20 to 20 N·m in inversion-eversion torquewere 0.026±0.002 s and 0.027±0.002 s.d. with 3.0% and 3.2% overshoot,respectively, as shown in FIG. 9D. With desired plantarflexion torqueoscillating between 10 and 90 N·m, the −3 dB magnitude and 450 phasemargin crossover frequencies were 27.2±0.2 Hz and 20.3±0.3 Hz,respectively (FIG. 8E). With desired inversion-eversion torqueoscillating between −20 and 20 N·m, the −3 dB magnitude and 450 phasemargin crossover frequencies were 29.8±0.2 Hz and 23.8±0.3 Hz,respectively, as shown in FIG. 8F.

FIG. 9 shows disturbance rejection, depicted as the relationship betweentorque error (% of the constant desired value) versus the frequency ofan applied disturbance in forefoot position. This characterizes theability of the system to reject unexpected environmental disturbances,such as from sudden contact with the ground. Torque error was less than30% of the desired value of 30 N·m for disturbance frequencies up to 18Hz.

Referring to FIG. 9, when a 0.012 m amplitude chirp disturbance wasapplied in forefoot endpoint position and commanded a constant desiredtorque of 30 N·m, torque error was less than 30% up to a disturbancefrequency of 18 Hz. This disturbance frequency and amplitude are similarto unexpected contact with stiff ground at a rate of 1.4 m/s.

FIG. 10 presents torque tracking during walking experiments. Desiredankle inversion torque was set to A Maximum, B 15 N·m, C zero, D −15N·m, and E Maximum Negative, while desired plantarflexion torque was aconsistent function of ankle plantarflexion angle. Maximum and MaximumNegative allowable inversion torque were limited by desiredplantarflexion torque, since forefoot ground reaction forces could notbecome negative. In each 100-stride trial, measured torque closelymatched desired torque, with RMS errors of at most 3.7 N·m inplantarflexion and 1.1 N·m in inversion-eversion across conditions.Differences between average torque and individual-step torques weredominated by changes in desired torque arising from natural variabilityin the subject's gait pattern.

Referring now to FIG. 10, during walking trials, subject 10 walkedcomfortably with the prosthesis while five levels of constant desiredinversion-eversion torque were applied. Torque tracking errors in bothplantarflexion and inversion-eversion directions were low across allconditions, with maximum RMS errors across the entire trial of 3.2 N·m(3.7% of peak) in plantarflexion torque and 1.1 N·m (3.8% of peak) ininversion-eversion torque, as shown in Table 2.

TABLE 2 Plantarflexion Torque Tracking Inversion-Eversion TorqueTracking Inversion- AVG AVG Eversion RMS RMS % % RMS % Torque Error %τ_(error) Error τ_(error) RMS Error τ_(error) Error τ_(error) τ_(inv) =3.2 ± 1.1 3.7% 1.3 N · m 1.0% 1.1 ± 0.4 3.8% 0.4 N · m 1.6% Maximum N ·m N · m τ_(inv) = −15 1.0 ± 0.4 2.2% 0.7 N · m 0.8% 0.9 ± 0.2 5.9% 0.7 N· m 4.4% N · m N · m N · m τ_(inv) = 0 2.9 ± 1.7 2.8% 0.6 N · m 0.6% 0.8± 0.2 — 0.5 N · m — N · m N · m τ_(inv) = 15 2.9 ± 0.8 2.8% 0.9 N · m0.9% 0.8 ± 0.2 5.6% 0.3 N · m 2.1% N · m N · m N · m τ_(inv) = Neg. 3.0± 0.9 3.3% 1.3 N · m 1.4% 1.0 ± 0.3 3.3% 0.4 N · m 1.6% Max. N · m N · m

In a second embodiment of the invention, the end-effector includes twotorque-controlled forefoot digits and a torque-controlled heel, whereinthe actuation of the three digits allows the adjustment of the originand magnitude of the resultant force vector. When the emulator has fullground contact, the origin of the vector can be approximately controlledwithin a support triangle, shown in FIG. 15, which has a maximum length(anterior-posterior) of 0.20 m and a maximum width (medio-lateral) of0.074 m. The peak force magnitude is approximately 2300 N. The emulatorutilizes vertical translation, sagittal rotation, and frontal rotationto maintain full ground contact for about 70% of stance time. Thisembodiment could be used in experiments investigating of load acceptancemechanics, manipulation of roll-over shape, and control of center ofpressure trajectory.

In the second embodiment, end-effector 300, shown in FIG. 11, iscentered around frame 302, Lateral forefoot 306 and a medial forefoot308, as well as heel 318 are joined to frame 302 via revolute joint 312.All three digits are independently controlled and sensed.

Each forefoot/heel (digit) acts as a lever wherein one end is pulled upby the motor-controlled cable, which results in a ground reaction forceat the other end. Lateral forefoot 306 is controlled by cable 412,having rope 314 a inside. Medial forefoot 308 is likewise controlled bycable 316 (rope not shown) and heel 318 is controlled by cable 322,having rope 322 a inside. The other side of cables 314, 316 and 322 areattached to motors 338, 340 and 342 respectively, shown in FIG. 12.Preferably, cables 314, 316 and 322 are Bowden cables consisting of asteel outer conduit and a Vectran inner rope.

Retraction springs 309, 310 and 320 connected respectively to lateralforefoot 306, medial forefoot 308 and heel 318, oppose the actuation totake up slack in the inner ropes.

Frame 302 consists of two mirrored frame halves joined by a connectingbridge. Joint torques are measured via strain gauge bridges 329 andjoint angles are measured via rotary encoders 328. Strain gauges 329, asshown in FIG. 11, are covered by a protective housing.

Heel 318 is preferably equipped with a limited-slip heel pad 32, whichallows the rubber on the bottom of the heel to slide smoothly instead ofscrubbing against the ground and wearing out. Additionally, each cable314, 316 and 322 is equipped with a strain relief 326, which preventsthe outer conduit from bending at sharp angles and fatiguing over time.

An off-the-shelf pyramid adapter 30, mounted ion the bridge of frame302, attaches end-effector 300 to the pylon on the user's existingprosthetic socket.

End-effector 300 is controlled by an off-board system, shown in FIG. 12,very similar to that of the first embodiment, shown in FIG. 2. Off boardmotors 338, 340 and 342 control digits 3076, 308 and 318 respectively onend-effector 300 under closed-loop computer control implemented assoftware running in computer 330. Tether 336 provides input signals fromencoder 328, and strain gauges 329, as well as any other sensors mountedon end-effector 300. Cables 334 convey signals to control the torquegenerated by motors 338, 340 and 342.

Using off-board motors 338, 340 and 342 and controller 330 allows a moreflexible, much higher-performance system with a simpler design and lessbody-mounted mass than an untethered system. In a preferred embodiment,controller 330 contains models of a plurality of commercially-availableprosthesis and is able to simulate the feel of these prostheses to thesubject.

In operation, the subject attached end-effector 300 to the user'sexisting prosthetic socket and the subject then walks on movingtreadmill 301. Feedback is collected via cable 336.

The dimensions of the end-effector were chosen to match the size of atypical human foot and the angles were chosen to exceed the range ofmotion found in normal walking. The heel-to-forefoot length of theend-effector was chosen to match the heel-to-metatarsal length ofatypical foot, similar to the design of commercial prosthetic feet.End-effector 300 was designed to be lightweight while still being ableto generate the forces equivalent to a 200 lb subject walking at 1.25m/s with a factor of safety of >1.5.

End-effector 300 has three active degrees of freedom, as shown in FIG.13(a-c), which can be expressed in terms of modes of movement orresultant force control. The modes of movement include verticaltranslation, shown in FIG. 14(a), sagittal rotation, shown in FIG. 14(b)and frontal rotation, shown in FIG. 14(c). Force control includesapproximate origin and magnitude control of the resultant force vector.The origin is located in the ground plane and is bounded by the supporttriangle of the emulator, as shown in FIG. 15.

Full control over the orthogonal component of the force produced by eachdigit can be controlled, but the total resultant force vector isaffected by both orthogonal and parallel components of all three digits.This uncertainty will lead to deviations in the expected origin,magnitude, and direction of the resultant force vector. However, thedesired and actual resultant force vectors will be similar if theparallel force components are small.

The weight of end-effector 300 is approximately 1.14 kg, not includingthe tethers. In terms of modes of movement, the vertical range oftranslation is approximately 0.050 m, the sagittal range of rotation isapproximately −32° to 26° counterclockwise from horizontal, and thefrontal range of rotation is approximately ±57° from vertical. In termsof force control capabilities, the maximum length and width of thesupport triangle is preferably 0.20 m and 0.074 m and the peak forcemagnitude is approximately 2300 N. Peak force magnitude decreases as thesupport triangle becomes smaller or if the force origin approaches theedge of the triangle. Resultant force magnitudes can also be expressedin terms of planar torques. Peak dorsiflexion, plantarflexion, andinversion-eversion torques are 68 Nm, 195 Nm, and ±24 Nm.

These results only include configurations where the emulator has allthree digits in contact with the ground. The emulator is expected toachieve full ground contact for about 70% of stance time.

The emulator described here as the second embodiment approximatelycontrols the location and magnitude of the resultant force over a longportion of stance phase. This control allows the emulator to investigateconcepts such as load acceptance mechanics. For example, this emulatorcould be used to vary the energy dissipated during the beginning ofstance phase.

This emulator is able to produce high forces while still maintaining aworn mass roughly equivalent to that of an average human foot and halfof that of commercially-available mobile powered prostheses with onecontrolled degree of freedom.

Although the invention is illustrated and described herein withreference to specific embodiments, the invention is not intended to belimited to the details shown of specific embodiments described. Inparticular, it will be realized by one of skill in the art that variousmodifications may be made in the details and implementation withoutdeparting from the invention.

We claim:
 1. A system for assessing the gait capacity of user lower legamputees, comprising: an end effector, attachable to a user, the endeffector having a plurality of digit portions rotatably attached to aplatform about a pivot point; a first motor corresponding to a firstdigit portion of the plurality of digit portions; a second motorcorresponding to a second digit portion of the plurality of digitportions; a first cable coupling the first motor to the first digitportion, the first cable configured to apply a first actuating force tothe first digit portion; a second cable coupling the second motor to thesecond digit portion, the second cable configured to apply a secondactuating force to the second digit portion; wherein the first motor isconfigured to provide the first actuating force independently from thesecond motor being configured to provide the second actuating force; acontroller, for controlling the first motor independently from thesecond motor; and one or more sensors, mounted on the end effector, theone or more sensors configured to send data regarding one or moreoperation parameters of the end effector to the controller.
 2. Thesystem of claim 1 further comprising one or more subject sensors,attached to the user, the subject sensors configured to send dataregarding one or more physical parameters of the subject to thecontroller.
 3. The system of claim 1 wherein the one or more sensors areconfigured to collect data regarding rotational angles of one or moredigit portions of the plurality of digit portions.
 4. The system ofclaim 2 wherein a subject sensor of the one or more subject sensors isconfigured to collect data regarding one or more of a respiration and ablood oxygen level of the user.
 5. The system of claim 1 wherein thecontroller is configured to store data comprising one or more models ofone or more prostheses and wherein the controller is configured tosimulate the operation of the one or more prostheses by controlling thefirst motor, the second motor, or both the first and second motors. 6.The system of claim 1 further comprising a treadmill configured tomeasure speed data representing a walking speed of the user, thetreadmill further configured to send the speed data to the controller.7. The system of claim 1 wherein the one or more sensors are configuredto measure a strain force on the first digit portion, the second digitportion, or both the first and second digit portions.
 8. A device forassessing the gait capacity of user lower leg amputees, comprising: aplatform supporting a mounting for rigidly fixing the device to the legof a user; a plurality of digit portions comprising a first digitportion and a second digit portion; one or more sensors configured forsensing an angle of rotation of the first digit portion the second digitportion, or both the first and second digit portions about a pivotpoint; a first motor corresponding to the first digit portion of theplurality of digit portions by a first linkage and configured to apply afirst tensioning force to the first digit portion by the first linkage;a second motor corresponding to the second digit portion of theplurality of digit portions by a second linkage and configured to applya second tensioning force to the second digit portion by the secondlinkage; a first spring member attached between the first digit portionand the platform, wherein the first spring member is tensioned when thefirst motor applies the first tensioning force to the first digitportion; and a second spring member attached between the second digitportion and the platform, wherein the second spring member is tensionedwhen the second motor applies the second tensioning force to the seconddigit portion.
 9. The device of claim 8 further comprising: a controllerconfigured to control the first motor, the second motor, or both thefirst and second motors, the controller configured to control firstmotor, the second motor, or both the first and second motors to vary theamount of work done by a user walking with the device; wherein thecontroller is configured to receive data from the one or more sensorsrepresenting a rotation of one or both of the first and second digitportions.
 10. The device of claim 9 wherein the controller is configuredto receive data representing one or more physical characteristics of theuser.
 11. The device of claim 10 wherein the one or more physicalcharacteristics includes a respiration measurement and a blood oxygenlevel of the user.
 12. The device of claim 9 wherein the controller isconfigured to store data comprising one or more models of prostheses andis further configured to cause one or both of the first and secondmotors to simulate one or more characteristics of the one or moreprostheses by varying the first tensioning force, the second tensioningforce, or both the first and second tensioning forces.
 13. A device forassessing the gait capacity of user lower leg amputees, comprising: aplatform supporting a mounting for rigidly fixing the device to the legof the user; a lateral toe portion, rotatably attached to the platformat a pivot point, the lateral toe portion having a first end forcontacting a walking surface and a second end extending past said pivotpoint; a lateral linkage, attached to the second end of the lateral toeportion; a medial toe portion, rotatably attached to the platform at apivot point, the medial toe portion having a first end for contacting awalking surface and a second end extending past the pivot point; amedial linkage, attached to the second end of the medial toe portion; aheel portion, rotatably attached to the platform at a pivot point, theheel portion having a first end for contacting a walking surface and asecond end extending past the pivot point; a heel linkage, attached tothe second end of the heel portion; a lateral motor, attached to thelateral linkage, for applying tension to the second end of the lateraltoe portion; a medial motor, attached to the medial linkage, forapplying tension to the second end of the medial toe portion; and a heelmotor, attached to the heel linkage, for applying tension to the secondend of the heel portion.
 14. The device of claim 13 further comprising:a lateral spring, attached at one end near the first end of the lateraltoe portion and at an opposite end to the platform; a medial spring,attached at one end near the first end of the medial toe portion and atan opposite end to the platform; and a heel spring, attached at one endnear the first end of the heel portion and at an opposite end to theplatform.
 15. The device of claim 13 further comprising: a lateralsensor, for sensing an angle of rotation of the lateral toe portionabout the pivot point; a medial sensor, for sensing an angle of rotationof the medial toe portion about the pivot point; and a heel sensor, forsensing an angle of rotation of the heel portion about the pivot point.16. The device of claim 15 further comprising: a controller forcontrolling the lateral motor, the medial motor and the heel motor, thecontroller being configured to individually control each of the motorsto vary the amount of work done by the user walking with the device;wherein the controller is configured to receive data from one or more ofthe lateral, medial, and heel sensors representing the respective anglesof rotation of the lateral toe portion, the medial toe portion and theheel portion.
 17. The device of claim 16 further comprising: a lateraltension encoder for sensing a tensioning force applied by the lateralmotor to the lateral toe portion; and a medial tension encoder forsensing a tensioning force applied by the medial motor to the medial toeportion; wherein the lateral tension encoder and the medial tensionencoder are each configured to send encoded tension data to thecontroller.
 18. The device of claim 17 wherein the controller isconfigured to store data comprising one or more models of prostheses andfurther wherein the controller is configured to cause one or more motorsto simulate one or more characteristics of the one or more prostheses byvarying tensioning forces applied to the lateral and the medial toeportions and the heel portion.
 19. The device of claim 17 furthercomprising a treadmill configured to measure speed data representing agait speed of the user, the treadmill configured to send the speed datato the controller.
 20. A system for assessing the gait capacity of userlower leg amputees, comprising: an end effector, attachable to a user,the end effector having a plurality of digit portions rotatably attachedto a platform about a pivot point; one or more motors corresponding toone or more respective digit portions of the plurality of digitportions, the one or more motors being located off-board relative to theend effector; one or more cables respectively coupling the one or moremotors to the one or more respective digit portions, the one or morecables configured to apply one or more respective actuating forces tothe one or more respective digit portions; a controller, for controllingthe one or more motors; and one or more sensors, mounted on the endeffector, the one or more sensors configured to send data regarding oneor more operation parameters of the end effector to the controller. 21.The system of claim 20 further comprising one or more subject sensors,attached to the user, the subject sensors configured to send dataregarding one or more physical parameters of the subject to thecontroller.
 22. The system of claim 21 wherein a subject sensor of theone or more subject sensors is configured to collect data regarding oneor more of a respiration and a blood oxygen level of the user.
 23. Thesystem of claim 20 wherein the one or more sensors are configured tocollect data regarding rotational angles of one or more digit portionsof the plurality of digit portions.
 24. The system of claim 20 whereinthe controller is configured to store data comprising one or more modelsof one or more prostheses and wherein the controller is configured tosimulate the operation of the one or more prostheses by controlling oneor more motors.
 25. The system of claim 20 further comprising atreadmill configured to measure speed data representing a walking speedof the user, the treadmill further configured to send the speed data tothe controller.
 26. The system of claim 20 wherein the one or moresensors are configured to measure a strain force on the one or morerespective digit portions of the plurality of digit portions.